This invention relates generally to methods and apparatus for computed tomographic cardiac imaging systems, and more particularly to methods and apparatus specialized for cardiac imaging with substantial component reuse.
In at least one known computed tomography (CT) imaging system configuration, an x-ray source projects a fan-shaped beam which is collimated to lie within an X-Y plane of a Cartesian coordinate system and generally referred to as the “imaging plane”. The x-ray beam passes through the object being imaged, such as a patient. The beam, after being attenuated by the object, impinges upon an array of radiation detectors. The intensity of the attenuated beam radiation received at the detector array is dependent upon the attenuation of the x-ray beam by the object. Each detector element of the array produces a separate electrical signal that is a measurement of the beam attenuation at the detector location. The attenuation measurements from all the detectors are acquired separately to produce a transmission profile.
In known third generation CT systems, the x-ray source and the detector array are rotated with a gantry within the imaging plane and around the object to be imaged so that the angle at which the x-ray beam intersects the object constantly changes. A group of x-ray attenuation measurements, i.e., projection data, from the detector array at one gantry angle is referred to as a “view”. A “scan” of the object comprises a set of views made at different gantry angles, or view angles, during one revolution of the x-ray source and detector. In an axial scan, the projection data is processed to construct an image that corresponds to a two dimensional slice taken through the object. One method for reconstructing an image from a set of projection data is referred to in the art as the filtered back projection technique. This process converts the attenuation measurements from a scan into integers called “CT numbers” or “Hounsfield units”, which are used to control the brightness of a corresponding pixel on a cathode ray tube display.
More particularly, and referring to FIGS. 5 and 6, a computed tomograph (CT) imaging system 10 is shown as including a gantry 12 representative of a “third generation” CT scanner. Gantry 12 has an x-ray source 14 that projects a beam of x-rays 16 toward a detector array 18 on the opposite side of gantry 12. Detector array 18 is formed by detector elements 20 which together sense the projected x-rays that pass through an object 22, for example a medical patient. Each detector element 20 produces an electrical signal that represents the intensity of an impinging x-ray beam and hence the attenuation of the beam as it passes through patient 22. During a scan to acquire x-ray projection data, gantry 12 and the components mounted thereon rotate about a center of rotation 24. Detector array 18 may be fabricated in a single slice or multi-slice configuration. In a multi-slice configuration, detector array 18 has a plurality of rows of detector elements 20, only one of which is shown in FIG. 2.
Rotation of gantry 12 and the operation of x-ray source 14 are governed by a control mechanism 26 of CT system 10. Control mechanism 26 includes an x-ray controller 28 that provides power and timing signals to x-ray source 14 and a gantry motor controller 30 that controls the rotational speed and position of gantry 12. A data acquisition system (DAS) 32 in control mechanism 26 samples analog data from detector elements 20 and converts the data to digital signals for subsequent processing. An image reconstructor 34 receives sampled and digitized x-ray data from DAS 32 and performs high speed image reconstruction. The reconstructed image is applied as an input to a computer 36 which stores the image in a mass storage device 38.
Computer 36 also receives commands and scanning parameters from an operator via console 40 that has a keyboard. An associated cathode ray tube display 42 allows the operator to observe the reconstructed image and other data from computer 36. The operator supplied commands and parameters are used by computer 36 to provide control signals and information to DAS 32, x-ray controller 28 and gantry motor controller 30. In addition, computer 36 operates a table motor controller 44 which controls a motorized table 46 to position patient 22 in gantry 12. Particularly, table 46 moves portions of patient 22 through gantry opening 48.
In a multislice imaging system 10, detector array 18 comprises a plurality of parallel detector rows, wherein each row comprises a plurality of individual detector elements 20. An imaging system 10 having a multislice detector 18 is capable of providing a plurality of images representative of a volume of object 22. Each image of the plurality of images corresponds to a separate “slice” of the volume. The “thickness” or aperture of the slice is dependent upon the thickness of the detector rows.
For example, and referring to FIGS. 7 and 8, a multislice detector array 18 includes a plurality of detector modules 50. Each detector module 50 has a plurality of detector elements 20. Particularly, each x-ray detector module 50 includes a plurality of photodiodes 52, a semiconductor device 54, and at least one flexible electrical cable 56. Scintillators 58, as known in the art, are positioned above and adjacent photodiodes 52. Photodiodes 52 may be individual photodiodes or a multidimensional photodiode array. Photodiodes 52 are optically coupled to scintillators 58 and generate electrical outputs on lines 60 representative of light generated by scintillators 58. Each photodiode 52 produces a separate electrical output 60 that is a measurement of the beam attenuation for a specific detector element 20. In one known embodiment, photodiode output lines 60 from each detector module 50 are located at the top and bottom of the photodiode array.
Semiconductor device 54, in one embodiment, includes two semiconductor switches 62 and 64. Switches 62 and 64 each include a plurality of field effect transistors (FETs) (not shown) arranged as a multidimensional array. Each FET includes an input line electrically connected to a photodiode output 60, an output line, and a control line (not shown). FET output and control lines are electrically connected by flexible cable 56. Particularly, one-half of photodiode output lines 60 are electrically connected to each FET input line of switch 62 with the remaining one-half of photodiode output lines 60 electrically connected to FET input lines of switch 64.
Flexible electrical cable 56 includes a first end (not shown), a second end (not shown) and a plurality of electrical wires 66 traveling therebetween. Cable 56 may, for example, be a single cable having multiple first ends 68 and 70 or in another known embodiment, may include multiple cables (not shown) each having a first end (not shown). FET output and control lines are electrically connected to cable 56 by wire bonding. Cable first ends 68 and 70 are secured to detector module 50 using mounting brackets 72 and 74. Detector modules 50 are secured to detector array 18 using rails 76 and 78.
One known detector array 18 is arcuate. However, and referring to FIG. 9, detector arrays are represented in simplified drawings by a flat, two-dimensional representation of the area exposed to radiation beam 16. In such representations, the axis of rotation of gantry 12 defines a z-direction of detector array 18. A transverse direction, i.e., the direction in which each row of detector elements 20 extends, defines an x-direction. In FIG. 9, rows (not separately shown) of detector elements 20 extend linearly in the plane of the paper, but each row, in reality, follows the arc of detector array 18. Centerline 80 on FIG. 9 represents an imaginary line of a radiation beam 16 passing through an axis of rotation of gantry 12. Detector array 18 is at least approximately symmetric about centerline 80, i.e., it is operationally insignificant if there is a slight asymmetry in the number of detector cells 20 on each side of centerline 80.
FIG. 9 is not drawn to scale. In addition, only a few detector modules 50 are represented in FIG. 9. In one known imaging system, fifty-seven detector modules 50, each having 16 rows of 16 elements, are assembled in detector array 18.
One problem with known imaging systems 10 is that they do not have detector arrays 18 that provide a sufficient number of rows of detector elements 20 to image a heart or other organ of patient 22 in a single revolution of radiation source 14 and detector array 18. Thus, known cardiac CT imaging methods require multiple revolutions and a substantial amount of time (relative to a cardiac cycle).
It would, in principle, be possible to image an entire heart in a single revolution using a larger detector array 18 that had a sufficient number of detector rows to capture attenuation data from all parts of the heart. A CT imaging system 10 having such a detector array 18 would provide the advantage of reducing a patient's total radiation dosage during a cardiac scan. However, to provide acceptable resolution for diagnostic purposes, detector array 18 would have to generate massive amounts of data from a large total number of detector elements. Providing a data acquisition system 32 capable of handling such a large amount of data would be costly.
It is known to selectively combine data from a plurality of adjacent detector rows (i.e., a “macro row”) to obtain images representative of slices of different selected thicknesses, which also reduces the amount of data that must be handled by data acquisition system 32 during a scan. If a detector array 18 large enough to image an entire heart during a single revolution were provided, rows could be combined to reduce the amount of data generated. Alternately, detector elements 20 could simply be made larger to provide increased coverage without providing massive amounts of data. However, either of these alternatives runs a significant risk of reducing resolution to unacceptable levels.
It would therefore be desirable to provide methods and apparatus to provide satisfactory CT cardiac imaging with a minimum number of revolutions of an x-ray source and detector. It would further be desirable if such imaging could be accomplished with a single revolution. It would also be desirable to reduce the amount of data collected during such a cardiac CT scan without making unacceptable sacrifices in image quality and resolution.